U.S. patent application number 11/947834 was filed with the patent office on 2016-07-14 for nanostructured devices for detecting and analyzing biomolecules.
This patent application is currently assigned to General Electric Company. The applicant listed for this patent is Rui Chen, Anthony John Murray, An-Ping Zhang. Invention is credited to Rui Chen, Anthony John Murray, An-Ping Zhang.
Application Number | 20160202254 11/947834 |
Document ID | / |
Family ID | 56367383 |
Filed Date | 2016-07-14 |
United States Patent
Application |
20160202254 |
Kind Code |
A1 |
Zhang; An-Ping ; et
al. |
July 14, 2016 |
NANOSTRUCTURED DEVICES FOR DETECTING AND ANALYZING BIOMOLECULES
Abstract
A biosensing FET device, comprising a plurality of
nanostructured SOI channels, that is adapted to operate in
solutions having a high ionic strength and provides improves
sensitivity and detection. Generally, the biosensing device
comprises an underlying substrate layer, an insulator and a
semiconductor layer and a plurality of channels in the
semiconductor layer comprising a plurality of whole or partially
formed nanopores in the channels.
Inventors: |
Zhang; An-Ping; (Rexford,
NY) ; Murray; Anthony John; (Lebanon, NJ) ;
Chen; Rui; (Clifton Park, NY) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Zhang; An-Ping
Murray; Anthony John
Chen; Rui |
Rexford
Lebanon
Clifton Park |
NY
NJ
NY |
US
US
US |
|
|
Assignee: |
General Electric Company
Schenectady
NY
|
Family ID: |
56367383 |
Appl. No.: |
11/947834 |
Filed: |
November 30, 2007 |
Current U.S.
Class: |
435/287.2 |
Current CPC
Class: |
G01N 27/4145 20130101;
G01N 33/54373 20130101; G01N 33/545 20130101; G01N 27/4146
20130101; B82Y 40/00 20130101; B82Y 15/00 20130101; G01N 27/327
20130101 |
International
Class: |
G01N 33/543 20060101
G01N033/543; G01N 27/414 20060101 G01N027/414 |
Claims
1. A biosensing FET device having a sensing surface, comprising: a
semiconductor layer comprising one or more channels having one or
more nanopores in the channel, wherein the nanopores comprise a
biomolecule; an insulator layer; an underlying substrate layer; and
wherein the biosensing FET device is adapted to operate in
solutions having an ionic strength that is equal to or less than
200 mM.
2. The biosensing device of claim 1, wherein one or more of said
channels has a height and one or more of said nanopores has a depth
that is less than the height of one or more of said channels
3. The biosensing device of claim 1, wherein one or more of the
channels has a height, and wherein one or more of the nanopores in
the channels has a depth that is the same as the height of one or
more of the channels.
4. The biosensing device of claim 3, wherein the nanopores have a
depth that is greater than the height of one or more of the
channels so that the nanopores extend through the top surface of,
and partially into, the insulator layer, and wherein the nanopores
have an inside surface, at least a portion of which is
functionalized, that is partially in the semiconductor layer and
partially in the insulator layer.
5. The biosensing device of claim 4, wherein the nanopores are
functionalized, at least in part, with one or more binders.
6. The biosensing device of claim 1, wherein one or more of the
nanopores is functionalized, at least in part, with one or more
binders.
7. The biosensing device of claim 1 wherein the nanopores in the
channel of the semiconductor layer locate one or more binding sites
proximate the sensing surface.
8. The biosensing device of claim 7, is adapted to operate in
solutions having an ionic strength that is equal to or less than 10
mM, and wherein the nanopores locate one or more of the binding
sites less than or equal to 5 nm from the sensing surface.
9. The biosensing device of claim 7, wherein the nanopores locate
the binding sites less than or equal to 1 nm from the sensing
surface.
10. The biosensing device of claim 1, wherein the channels have a
density of nanopores that is between 10.sup.10 to 10.sup.12 per
cm.sup.-2.
11. The biosensing device of claim 1, wherein the semiconductor
layer comprises silicon.
12. A biodetector comprising the biosensing device of claim 1.
13-23. (canceled)
Description
BACKGROUND
[0001] The invention relates generally to nanostructured devices
for detecting or analyzing biomolecules and their interactions.
[0002] Proteomics offers great potential for discovering biomarker
patterns for earlier screening and detection of lethal and
infectious diseases, systematic monitoring of physiological
responses to drugs, and selecting the best treatment options for
individual patients. For routine clinical use, an inexpensive,
easy-to-use, multiplexed and high throughput protein analysis
platform is needed, with high sensitivity and specificity for
detection of low-abundance biomarkers in serum or other body
fluids. There is also a need for high throughput and highly
integrated sensor arrays for drug screening.
[0003] Nanostructured sensor arrays that use purely electrical
detection, such as a field effect transistor (FET), fabricated with
Si or other semiconductors, offer some of the desired
characteristics. In such a device, a device channel of Si or other
semiconductors is defined between two electrodes. The surface of
the semiconductor channel or its oxide surface may be modified and
covalently functionalized with antibodies or other receptor ligands
for quantitative biorecognition. The binding of protein or other
biomolecules induces net charge change, or change in dipole moment
and binding-induced dipoles or modification of energy distribution
and/or density of surface states. These binding events can change
surface potential of the FET device and therefore modulate the
conductance of the semiconductor channel. A small voltage or
current, small enough not to disturb biomolecule interactions, is
applied between two electrodes, and the change in conductance of
the device channel is related and calibrated to the analyte
concentration in a solution. When the device channel is reduced to
nanoscale, the detection limit can be significantly reduced due to
increased surface-to-volume ratio. Further, the response time can
also be reduced due to favorable mass transport at low analyte
concentrations due to small binding capacity of the small sensing
surface. The ultralow detection limit of the nano-FET sensor at low
ionic strength solutions has been recently demonstrated.
[0004] However, these devices may be rendered ineffective due to
the screening effect in higher ionic strength solutions. The Debye
screening length is defined as the distance from the sensing
surface where potential change can be detected by the sensing
device. In a high ionic strength solution, the screening length is
reduced by ions and thus, analytes present beyond the screening
length cannot be detected. As shown in FIG. 1, the Debye screening
length decreases with an increase in ionic strength, therefore the
binding events may not be detectable in high ionic strength
solutions. It would be desirable to provide a method and a device
that would enable a nano-FET biosensor to operate at higher ionic
strengths when physiological samples with high ionic strength are
to be analyzed, such as analyzing protein biomarkers in serum or
other body fluids.
[0005] The Debye-Huckel Theory is useful to better understand the
issues associated with operating biosensing devices in higher ionic
solutions. For example, assuming a perfect orientation of an
immobilized antibody, FIG. 2 shows the interaction between an
antigen and an antibody in a solution. The potential distribution
as a function of distance away from the electrolyte-immuno FET
interface with immobilized antibodies is shown for both high ionic
strength and low ionic strength cases. It can be seen from FIG. 2
that the electrostatic potential decreases rapidly as the binding
site moves away from the electrolyte/gate insulator interface. The
Debye screening length, .delta., can be simply defined, in this
example, as the distance away from the electrolyte/gate insulator
interface at which a charge redistribution can still be detected by
the FET sensor. In the high ionic strength environment, the Debye
length is extremely short due to charge screening of the analyte
antigen by excess ions (or more precisely "counterions") present in
solution. From the FET perspective, this charge screening effect
makes it ineffective to detect charges induced by antigen/antibody
interaction beyond the screening distance and therefore the
anitigen molecule must come closer to the sensor surface in order
for its intrinsic or induced charges to be detected. Beginning with
a buffer solution with an ionic strength solution of 0.2M, for
example, the calculated Debye length is approximately lnm, which is
significantly shorter than the average length of an antibody
molecule (.about.10 nm). Therefore the binding of antigen to the
antibody receptor results in the redistribution of charges too
distant to be detected by the FET sensor. However, in the absence
of such excess charged species in solution, as in the case for low
ionic strength situation, the screening effect by counterions is
not as severe. The Debye length is much longer and the antigen
molecule can be detected at a distance that is further away from
the sensing surface. The overlapping of potentials in FIG. 2 in the
low ionic strength case signifies a measurable effect with
potentiometrically-operated immunoFET. The equation for Debye
screening length in electrolytic solution is illustrated as:
.delta. = K B T 8 .pi. e 2 I ##EQU00001##
where K.sub.B is the Boltzmann constant, T is temperature, e is the
elementary charge (1.6.times.10.sup.-19 C), e is the dielectric
constant, and I is the ionic strength which has the expression
I = 1 i n i Z i 2 2 ##EQU00002##
[0006] where ni represents the concentration of the ith ionic
species in the electrolytic solution and Zi is the charge of the
ith species. The sum of the product of the concentration and charge
of all ionic species gives an estimate of the ionic strength of the
electrolytic solution. Since the Debye length varies as the inverse
square root of the ionic strength, the sensing response depends on
the ionic strength of the solution.
[0007] The nanoscale channel can increase surface-to-volume ratio
of the device and therefore significantly lower the detection
limit, but lithography tools that are expensive and lower
throughput are required to define nanoscale patterns. It would also
be desirable to increase surface-to-volume ratio of the channel
without reducing the channel to nanoscale size, so larger channel
size can be used to achieve the low detection limit and more
conventional and inexpensive lithography tools can be used. It can
significantly reduce the cost of device fabrication.
BRIEF DESCRIPTION
[0008] The invention generally relates to a semiconductor sensing
device having a raised structure, referred to as device channel,
wherein the device channel comprises one or more nanopores whole or
partially formed in the raised structure on an underlying
insulating layer on a substrate. The invention also generally
relates to methods of making and using the sensing device that
comprises nanopores formed in a silicon-on-insulator structure
(SOI), such as a Si channel with nanopores. This nanopore structure
physically brings binding sites of antibodies or other receptor
molecules proximate to a sensing surface and enables the biosensor
to operate at a higher ionic strength. In addition, the nanopores
in the channels of the sensing device increase surface/volume ratio
of the device, and enable lower detection limits and greater
sensitivity at larger device channel size.
[0009] The devices and methods use nanostructured SOI channels to
enable the sensor to operate in higher ionic strength solutions. In
one embodiment, nanopores are formed in the device channel and stop
at the underlying SiO2 layer. The antibodies or other receptor
ligands can be selectively functionalized on the underlying SiO2
layer inside the nanopores. In another embodiment, a thin metal
layer, such as Au or Ag, may be selectively deposited on the
underlying SiO2 layer inside the nanopores by a lift-off process.
Antibody or other receptor ligands can be selectively
functionalized on the metal surface. The nanopores effectively
locate the binding sites proximate to the sensing surface. The size
and pitch of the nanopores may be controlled using block copolymer
methods or other suitably controllable nanopatterning methods. In
another embodiment the device comprises an underlying substrate
layer, an insulator and a semiconductor layer and the one or more
channels on the underlying insulator layer comprising one or more
nanopores of varying depths in the channels.
[0010] An embodiment of the biosensing FET device, of the
invention, having a sensing surface, generally comprises: a
semiconductor layer comprising one or more channels having one or
more nanopores in the channel; an insulator layer; and an
underlying substrate layer.
[0011] In one or more of the embodiments, the channels may have a
height and one or more of the nanopores has a depth in the
channels, that is less than, equal to, and greater than, the height
of one or more of the channels.
[0012] One or more of the embodiments may comprise nanopores that
are functionalized. For example, the nanopores may have a depth
that is greater than the height of one or more of the channels so
that the nanopores extend through the top surface of, and partially
into, the insulator layer, and wherein the nanopores have an inside
surface, at least a portion of which is functionalized, that is
partially in the semiconductor layer and partially in the insulator
layer. Although not intended to be limiting, the all or part of the
surface of the nanopores may functionalized, for example, with one
or more binders. The nanopores in the semiconductor channel may be
adapted to locate one or more binding sites proximate the sensing
surface. As a non-limiting example, the device may be adapted to
operate in solutions having an ionic strength that is equal to or
less than 10 mM; wherein the nanopores are adapted to locate one or
more of the binding sites less than or equal to 5 nm from the
sensing surface. As another example, the device may be adapted to
operate in solutions having an ionic strength that is equal to or
less than 200 mM; and wherein the nanopores locate the binding
sites less than or equal to 1 nm from the sensing surface. The
channels may have a range of densities of nanopores such, as but
not limited to, between 10.sup.10 to 10.sup.12 per cm.sup.-2.
[0013] Any one or more of the embodiments of the biosensing device
may be incorporated into a biosensing detector.
[0014] An embodiment of the method of the invention, of making a
biosensing device, generally comprises the steps of: a) providing
an underlying substrate layer; b) disposing an insulator on the
substrate; c) disposing a semiconductor, having an exposed surface
with one or more channels, on the insulator layer; and d) forming
one or more nanopores in one or more of the channels; wherein the
channels may have a density of nanopores between 10.sup.10 to
10.sup.12 per cm.sup.-2. The nanopores may also be
functionalized.
[0015] As a non-limiting example, the nanopores may be formed in
the channels to achieve a density of nanopores between
4.times.10.sup.10 to 2.times.10.sup.11 per cm.sup.-2 and wherein
the nanopores have a pitch between 20 nm to 50 nm. One or more of
the channels has a height and one or more of the nanopores has a
depth that is less than, equal to, or greater than, the height of
one or more of the channels.
[0016] The nanopores may be formed by, but not limited to,
nanopatterning, such as, block copolymer lithography; wherein block
copolymer lithography may comprise the steps of: (a) coating the
semiconductor with a block copolymer capable of phase separating;
(b) providing stimulus to form phase separated block copolymer; (c)
etching the semiconductor layer to form one or more nanopores; and
(d) removing at least a portion of the phase separated block
copolymer. The block copolymer may comprise, but is not limited to,
one or both of polystyrene-block-polybutadiene and
polystyrene-block-polyisoprene.
[0017] One or more of the channels may have a height and one or
more of the nanopores may have a depth that is equal to or greater
than the height of one or more of the channels, so that the
nanopores extend through the semiconductor layer and into the
insulator layer; and wherein the nanopores may be silanized.
DRAWINGS
[0018] These and other features, aspects, and advantages of the
present invention will become better understood when the following
detailed description is read with reference to the accompanying
drawings in which like characters represent like parts throughout
the drawings, wherein:
[0019] FIG. 1 is a graph illustrating a decrease in the Debye
Screening Length as the ionic strength of an ionic solution
increases,thus disabling a biosensing device from operating in a
high ionic solution;
[0020] FIG. 2 is a graph illustrating the influence of solution
ionic strength on the response of nano-FET sensor;
[0021] FIG. 3 illustrates an embodiment of a channel of a
biosensing device in which the nanopores are formed by etching
completely through the semiconductor layer;
[0022] FIG. 4 illustrates an embodiment of a channel of a
biosensing device in which the nanopores are partially etched
through the semiconductor layer to form the nanopores having a
depth that is less than the height of the channel;
[0023] FIGS. 5A and 5B illustrate a channel with nanopores located
therein having diameters of approximately 10-20 nm;
[0024] FIG. 6 is a diagram of Surface/Volume ratio vs. Channel
length for channels shown in FIGS. 5A and 5B, with and without
nanopores where the channel width is approximately 20 nm or 50
nm;
[0025] FIGS. 7A-7B illustrate embodiments of a channel of a
substrate with partial nanopores located therein;
[0026] FIG. 8 is a diagram of the Surface/Volume ratio vs. Channel
Length for channels shown in FIGS. 7A and 7B, with and without
partial nanopores, where the channel width is approximately 20 nm
or 50 nm.
[0027] FIG. 9 is a photograph of an embodiment of the nanopores
with a diameter that is approximately 15 nm; and
[0028] FIGS. 10A to 10I illustrate embodiments of the steps for
fabricating the biosensing device of the invention.
DETAILED DESCRIPTION
[0029] The present disclosure provides an embodiment of a
biosensing device in which nanopores are etched into the channels
of a nano-field effect transistor (FET) device. The nanopores are
formed as described below by block copolymer nanolithography or by
other nanopatterning techniques such as nanoimprint, soft
lithography, or by e-beam lithography, etc. as described below.
[0030] Referring to the drawings, FIG. 3 shows a biosensing device
10 as one embodiment of the invention. The biosensing device
comprises a substrate layer 12. The substrate layer may be made
from a material such as silicon or glass. The biosensing device
further comprises an insulator layer 14 on the top surface of the
substrate. The insulator layer may be made from an insulating
material, such as but not limited to, silicon oxide, silicon
nitride, and the like. On the surface of the insulator layer 14
formed on the substrate layer 12, a semiconductor layer 16 is made
available in the form of raised structures characterized by height,
to form device channels with nanopores 18 on the insulator layer.
An exemplary material used to make semiconductor layer includes
silicon. The height of the semiconductor layer is generally uniform
in a given biosensing device, and ranges from about lnm to about
1000 nm preferably 5 nm to 20 nm. Within the channels, nanopores 18
are present. The underlying insulating layer exposed inside
nanopores, such as silicon oxide, is functionalized with binders,
such as, but not limited to, antibodies or other receptor molecules
or receptor ligands to provide binding sites, which are used for
identification and quantifying of analytes. As used herein, a
binder generally refers to molecules that have binding affinities
either with themselves or more commonly with other molecules.
Generally, the binder specifically binds to an analyte of interest.
For example, such molecules include, but are not limited to,
antigens, antibodies, affibodies, nanobodies, an enzyme, an enzyme
substrate/inhibitor, aptamer and nucleic acids. This list is not
exhaustive. Such binders may be used to functionalize one or more
surfaces of the devices.
[0031] The biosensing device further comprises a source electrode
and drain electrode. The biosensing device is useful in identifying
and quantifying analytes such as, but not limited to, antigens,
antibodies, nanobodies, affibodies, aptamers, nucleic acids,
proteins, viruses and other chemical moieties.
[0032] The presence of nanopores in the device channels locates the
binding sites close to the sensing surface of the device. As a
result, the charges, induced by the binding events at the binding
sites, may be detected by the biosensing device, even in the
presence of high ionic strength solutions. The presence of
nanopores in the device channels also increases the
surface-to-volume ratio of the device channel. This results in
increased sensitivity towards the analytes to be detected. Thus,
the biosensing device is useful for identifying and quantifying
analytes in solution, including solutions having high ionic
strength. The biosensing device reduces or eliminates the need for
extensive desalting steps of the solution containing analytes
before the detection. Further, with low detection limits, simple
dilution by low ionic strength buffer can lower the ionic strength
to a level allowing detection by the biosensing device of the
invention. The presence of nanopores in the device channel also
increases surface-to-volume ratio and therefore achieve very low
detection limit even at large channel width. Thus expensive
lithography steps can be avoided. As shown in FIG. 1, it can be
seen that, for antibodies having an average size of approximately
10 nm, the nanowire sensor without nanopores can detect binding
events in approximately 1 mM ionic solutions. If the presence of
nanopores brings the binding sites to approximately 5 nm away from
the sensing surface, the sensor can work in an approximately 10 mM
ionic strength solutions. Similarly, with a 1 nm separation between
binding sites and the sensing surface with nanopores, the nanowire
sensor can operate in 200 mM ionic strength solutions.
[0033] FIG. 4 shows a biosensing device 22 with incompletely etched
nanopores 18 in the semiconductor channel The biosensing device
comprises a substrate layer 12. The biosensing device further
comprises an insulator layer 14 on the top surface of the
substrate. On the surface of the insulator layer 14 formed on the
substrate layer 12, a semiconductor layer 20 is made available in
the form of raised structures characterized by height, to form
device channels with nanopores 18 in the semiconductor layer. An
exemplary material used to make semiconductor layer includes
silicon. The height of the semiconductor layer is generally uniform
in a given biosensing device, and ranges from about lnm to about
1000 nm preferably 5 nm to 20 nm. The nanopores 18 is incompletely
etched in semiconductor layer 20. The presence of nanopores 18 can
increase surface-to-volume ratio of the semiconductor channel and
enhance single-to-noise ratio of the sensing device. Therefore the
detection limit can be lowered and sensitivity of the biosensing
device is improved. The exposed sensing surface of the
semiconductor channel is functionalized with antibodies or other
receptor molecules or receptor ligands to provide binding sites,
which are used for identification and quantifying of analytes.
[0034] The nanopores in the device channel increase the
surface-to-volume ratio and sensitivity at larger device channel
width. FIG. 6 is a graph of the embodiments shown in FIGS. 5A and
5B illustrating the difference between channels with 20 nm
thickness and channels with 50 nm thickness with and without
nanopores. The diameter of the nanopores is 10 nm and the pitch is
20 nm. As seen in line 51, a channel with a 50 nm thickness without
nanopores decreases in surface/volume ratio from 60 .mu.m.sup.-1
about 22 .mu.m.sup.-1 when the channel length increases from 50 nm
to 1000 nm. While the same channel, as shown in line 50, with the
same thickness of 50 nm channels with nanopores maintains
surface/volume ratio of approximately 50 as the channel length
increases from about 50 nm to 1000 nm.
[0035] Similarly FIG. 6 shows, in line 43, that a channel with 20
nm thickness without nanopores that the surface to volume ratio
drops from about 90 .mu.m.sup.-1 to at a channel width of about 50
nm to a surface to volume ratio of about 52 .mu.m.sup.-1 for a
channel width of up to 1000 nm. Comparatively, as shown in line 42,
a channel with the same 20 nm length with nanosized pores remains
at about 100 .mu.m.sup.-1 for its surface to volume ratio as the
channel length increase from about 50 nm to up to 1000 nm. This
illustrates that completely etched nanopores in the semiconductor
channel increase surface-to-volume ratio and achieve high
sensitivity for larger channel width. The density of nanopores is
about 2.5.times.10.sup.11 cm.sup.-2 for 10 nm diameter nanopores
with 20 nm pitch, which is comparable to the antibody density
typically achieved on a Si/SiO2 surface (10.sup.10.about.10.sup.12
cm.sup.-2). Preferably the density is in the range of
4.times.10.sup.10 to 2.times.10.sup.11 cm.sup.-2 corresponding to
pitches of nanopores between 20 nm to 50 nm.
[0036] FIGS. 7A and 7B illustrate embodiments of the present
disclosure with partially etched nanopores fabricated on a channel
of an underlying substrate layer. The nanopore diameter is
approximately 10-20 nm. The channel width is greater or equal to 50
nm with a channel thickness of greater than or equal to 10 nm.
[0037] FIG. 8 is a graph of the embodiments shown in FIGS. 7A and
7B illustrating the difference between channels with 20 nm
thickness and channels with 50 nm thicknesses with and without
nanopores. As seen in line 47 a channel with a 50 nm thickness
without nanopores decreases in surface/volume ratio from 60
.mu.m.sup.-1 about 22 .mu.m.sup.-1 while, as seen in line 48, the
same channel with the same thickness of 50 nm channels with
nanopores has surface/volume ratio of approximately 170
.mu.m.sup.-1 at 100 nm channel length that decreases only to about
120 .mu.m.sup.-1 for channel length of nearly 1000 nm.
[0038] Similarly FIG. 8 shows, in line 52, that a channel with 20
nm thickness without nanopores that the surface-to-volume ratio
drops from about 90 .mu.m.sup.-1 at a channel width of about 50 nm
to a surface-to-volume ratio of about 52 .mu.m.sup.-1 for a channel
width of about 1000 nm. Comparatively, as seen in line 53, a
channel with the same 20 nm thickness with nanopores decreases
surface-to-volume from about 200 .mu.m.sup.-1 to 160 .mu.m.sup.-1
for a channel length of 1000 nm. This illustrates that the present
disclosure provides that partially formed nanopores on the
substrate layer also provide for increased surface-to-volume ratio
and achieve high sensitivity for larger channel length.
[0039] FIG. 9 is a photograph showing nanopores in a Si layer on a
substrate in accordance with the present disclosure having a
diameter of about 15 nm. The scale shown is 200 nm.
[0040] In another aspect, the present disclosure provides a method
for making a biosensing device. The method comprises providing a
substrate layer, an insulating layer and a semiconducting layer.
FIG. 10A-10I illustrate the fabrication of the device in accordance
with the teachings of the present disclosure. The device
fabrication starts with a silicon-on-insulator (SOI) structure 64
(FIG. 10A). The SOI wafer with different top Si layer thickness may
be purchased from a commercial source (SOITEC in France, for
example). Optionally, the top Si layer 65 may be thinned by thermal
oxidation and followed by a wet etch (buffered oxide etch) to
remove converted SiO2. The nanopores in device channel (top Si
layer) may be formed by block copolymer nanolithography, or by
other nanopatterning techniques (nanoimprint, soft lithography, or
e-beam lithography, etc.).
[0041] The block copolymer (BCP) 68 may be spun on the SOI wafer 64
(FIG. 10B). Exemplary block copolymers useful in the invention
include, but not limited to, polystyrene-polybutadiene (PS-PB),
postyrene-polyisoprene (PS-PB), polystyrene-b-poly(methyl
methacrylate) (PS-b-PMMA), and the like. Block copolymers are
composed of two different polymer chains covalently bonded together
on one end. Polymers are usually immiscible with one another and
phase-separate; in block copolymers, molecular connectivity forces
phase separation to occur on molecular-length scales. As a result,
periodically ordered nanometer-sized microdomains (such as
cylinders or spheres) form, and their specific chemical,
electrical, optical, or mechanical properties can be controlled by
the choice of the constituent polymers. The sizes and periods of
these microdomain structures are governed by the chain dimensions
and are typically on the order of 10-30 nm. Structures smaller than
10 nm are also obtainable if one chooses appropriate blocks with a
high Flory-Huggins interaction parameter and decreases the block
lengths.
[0042] For example, asymmetric polystyrene-polybutadiene (PS-PB)
diblock copolymer in toluene solution can be spin-coated onto the
SOI wafer and film thickness is controlled by varying spinning
speed and polymer concentration (FIG. 10C). In bulk, the PS-PI
separates into a spherical morphology and produces 70 (PI spheres)
in 69 (PS matrix) with body-centered-cubic order. The films are
then annealed at 125.degree. C., a temperature above their glass
transition temperatures, for 24 hours in vacuum to obtain
well-ordered morphologies. The microdomain monolayer film was
exposed to ozone to selectively degrade and remove the PB spherical
domains before a CF.sub.4 reactive ion etch (RIE) or
CF.sub.4/O.sub.2 RIE. Ozone predominantly attacks the carbon-carbon
double bonds in the PB backbone, cutting the bonds and producing PB
fragments that can be dispersed in water. This results in regular
spherical voids in the PS matrix and hence in a variation of the
effective total thickness of the copolymer mask. The regions
underneath the empty spheres are exposed to the RIE to produce
holes in silicon, whereas the rest is still protected.
Chlorine-based or fluorine-based reactive ion etching (RIE) is used
to etch Si, with the copolymer film itself as the etching mask
(FIG. 10D). Etching may be effected by several other techniques
known in the art, such as electron cyclotron resonance (ECR) high
density plasma etch or inductively coupled plasma (ICP) etch,
chemically assisted ion beam etching (CAIBE), wet chemical etch,
and the like. The depth of this Si etch may be controlled by
controlling etch time with a fixed etch rate, to achieve either
complete etch or incomplete etch.
[0043] After Si is completely removed from the nanopores and the
underlying SiO2 is exposed, an optional thin layer of thin metal,
such as Au or Ag, may be deposited on SiO2 surface 66 in the
nanopores by e-beam evaporation, thermal evaporation, sputtering,
or other metal deposition techniques. The remaining BCP may be
removed by oxygen plasma, or by solvents (acetone, etc.), or other
strippers (FIG. 10-E). The device active region (device channel) is
patterned by conventional photolithography for channel width
greater than 300 nm or nanopattering techniques for channel width
less that 300 nm, and then etched by a plasma etch or wet chemical
etch (FIG. 10F).
[0044] For the complete etch where Si is completely removed inside
nanopores, a silane layer can be selectively formed on bottom SiO2
but not on Si surface or on Au surface but not on Si or SiO2
surface FIG. 10G), and antibody or other receptor ligands can be
immobilized on silanized surface (FIG. 10H). In another embodiment,
a Si--C functionalization may be formed on Si surface and then
antibody or other receptor ligands is attached. For the incomplete
etch where Si is partially removed inside nanopores, the surface
can be modified and functionalized with antibodies or other
receptor ligands. This arrangement physically moves the binding
sites close to the sensing surface. Proteins or other biomolecules
complementary to the receptor ligands can then bind on the receptor
ligands and thus change conduction of the device channel (FIG.
10I).
[0045] The density of nanopores can be controlled by manipulating
the composition of the block copolymers. For example, density of
nanopores can be controlled between 10.sup.10.about.10.sup.12
cm.sup.-2, or preferably 4.times.10.sup.10 to 2.times.10.sup.11
cm.sup.-2 corresponding to pitches of nanopores between 20 nm to 50
nm. The presence of nanopores increases surface-to-volume of the
device channel, and therefore increase signal-to-noise ratio when
binding events modulate conductance of the device channel.
[0046] The biosensing device measures the variation in its
conductance due to the variation of the surface potential. In one
embodiment, a reference device without antibody or other receptor
molecules may be positioned close to the sensing device. The
response from the reference device may be subtracted from the
sensing device to account for non-specific binding. Other
components that may be used in the biosensing may include membranes
to filter particulate matter, a buffer solution, and so on.
[0047] In another aspect, the present disclosure provides a method
for analyzing analytes in solution using the biosensing device
described herein. The present disclosure also provides kits that
comprise biosensing device described herein.
[0048] While only certain features of the invention have been
illustrated and described herein, many modifications and changes
will occur to those skilled in the art. It is, therefore, to be
understood that the appended claims are intended to cover all such
modifications and changes as fall within the true spirit of the
invention.
* * * * *